From the earliest days of modern medicine the assessment of circulatory system viability has been of foremost concern to clinicians. If elements of the circulatory system become diseased the continued health of the tissues they supply is threatened. All manners of methods have been applied to the assessment of circulatory flow from non-invasive (requiring no physical insertion of instrument, a material or ionizing radiation) to highly invasive. Traditional non-invasive methods include plethysmography and ultrasound Doppler, see Mozersky et al, Ultrasonic Arteriography, Arch. Surg. Vol. 103, pp 663-667, Dec. 1971, while invasive methods range from thermodilution to the placement of electromagnetic flowmeters around individual blood vessels.
In recent years ultrasound has experienced considerable growth in clinical use. This growth is due in part to the development of color flow Doppler (see Eyer et al, Color Digital Echo/Doppler Image Presentation, Ultrasound in Med. & Biol., Vol. 7, pp 21-33 and Fox, Multiple Crossed-Beam Ultrasound Doppler Velocimetry, IEEE Transactions on Sonics and Ultrasonics, Vol. SU-25, pp 281-286), so called because the direction of flow in the body is displayed as a color superimposed on the tomographic display. In addition, ultrasound imaging has become widely accepted by the medical community because of its safety, ease of use and low cost. Color flow Doppler imaging provides a means of assessing blood flow in a two-dimensional tomograph, or slice through the body. This non-invasive, multi-dimensional, real-time measurement of blood velocity in the body provides a powerful clinical diagnostic tool. Despite the success of color flow Doppler processors, they suffer from certain major fundamental limitations. Color flow Doppler processors measure the axial component (along the beam) of the true velocity vector by detecting the frequency shift of the ultrasound transmit burst. Color flow Doppler processors do not measure the lateral or elevational vectors (across the beam). This approach to the analysis of flow fails to provide accurate perfusion rates because all three component vectors of the blood velocity must be known to calculate the flow from the velocity. Color flow Doppler processors frequently employ algorithms that impose a set of simplifying assumptions to estimate the actual flow present in the lumen, but they fall well-short of being able to provide accurate perfusion measurements in many common diagnostic situations.
In ultrasound velocity measurement, an ultrasound interrogation is performed by the transmission of an acoustic mechanical pulse into a test region or resolution volume A single interrogation consists of the transmission of a pulse of energy into the test region and the return of that energy from the test region. The electrical signal received from a pulse or series of pulses which interrogate a given area of the test region is called a line. Thus dividing the test region into one or more lines of interrogation. Interrogation is achieved in ultrasound by a transducer, excited by an electric pulse. The transducer converts the electrical energy into mechanical energy. The resolution volume is insonified, i.e. the mechanical energy is transmitted into the resolution volume. When the mechanical pulse impacts on an impedance boundary, some of the energy is reflected, some is transmitted and some is lost to absorption and scattering. Objects presenting an impedance boundary which scatters the incident wave are known as scatterers. The time between transmit and receive for a given target range is equal to C/2Z where C is the velocity of the sound in the body and Z is the target range. The factor of 2 results from the round-trip travel of the pulse to the target and back. When the reflected signal arrives back at the transducer the mechanical energy is converted back into electrical energy. The electrical signal may then be used for display purposes.
When a resolution volume containing a large number of scatterers is insonified, the resultant returned wave is a combination of constructive and destructive interference caused by the scatterers. This phenomenon was first described in the field of laser optics when it was observed that the light returning from an illuminated screen does not produce a uniform spot of reflected light. Instead a granular intensity variation, bearing no clear spatial relationship to the structure of the screen. As each scatterer reflects the incident waveform it modifies the amplitude and phase. The perceived signal is then the sum of all the individual reflected wavelets within a resolution volume. This summation process results in both constructive and destructive interference producing a random brightness pattern. Changes in this pattern result from changes in the distribution of scatterers in the illuminated field. If small changes in the distribution occur the pattern remains well correlated with the original pattern. The effect has been shown to occur in a similar fashion in ultrasound.
Considerable interest has been shown in tracking variations in amplitude of the resultant returned wave as it moves through the image with the motion of tissue and blood (see Embree et al, The Accurate Measurement of the Volume Flow of Blood by Time Domain Correlation, 1985 Ultrasonics Symposium, pp 963-966 and Bonnefous, Measurement of the Complete (3D) Velocity Vector of Blood Flows, 1988 Ultrasonics Symposium, pp 795-799). The primary attraction of this method of tracking is that it is inherently angle independent, in contrast to Doppler methods that require a component of the velocity to exist in the axial dimension. Work to date has focused on correlation processors in which a region of space is sampled in one acquisition and then tracked within the surrounding area by means of a correlation search process (see Trahey et al, Angle Independent Ultrasonic Detection of Blood Flows, IEEE Transactions on Biomedical Engineering, Vol. BME-34, pp 965-967 and Trahey et al, Angle Independent Ultrasonic Blood Flow Detection Frame-to-Frame correlation of B-mode Images, Ultrasonics 1988, Vol. 26, pp 271-276). The method has been proven in two dimensions using off-line simulations and has been implemented in a real time in one-dimensional processor (Philips Platinum CVI). Although correlation processors have been criticized because of the considerable computational complexity they have a more serious fundamental limitation. Correlation processors require kernel cells that must be tracked within the search region. In order for the correlation process to track successfully the form of the kernel must be unique among the closely surrounding regions. If the kernel is not unique the correlator will track to the wrong location. The need for a unique kernel dictates that the kernel must be at least two resolution cells long in any dimension tracking is to be performed. As the kernel cell size is increased, the imaging resolution decreases. Kernel cell size also presents particular problems at vessel walls where the kernel may include both stationary and moving amplitude variations. This problem is exacerbated by the need to use a kernel greater than the resolution cell size.